Magnetic resonance imaging (MRI) is a medical imaging modality that can create pictures of the inside of a human body without using x-rays or other ionizing radiation. MRI uses a powerful magnet to create a strong, uniform, static magnetic field (i.e., the “main magnetic field”). When a human body, or part of a human body, is placed in the main magnetic field, the nuclear spins that are associated with the hydrogen nuclei in tissue water become polarized. This means that the magnetic moments that are associated with these spins become preferentially aligned along the direction of the main magnetic field, resulting in a small net tissue magnetization along that axis (the “z axis,” by convention). A MRI system also comprises components called gradient coils that produce smaller amplitude, spatially varying magnetic fields when current is applied to them. Typically, gradient coils are designed to produce a magnetic field component that is aligned along the z axis (i.e., the “longitudinal axis”) and that varies linearly in amplitude with position along one of the x, y or z axes. The effect of a gradient coil is to create a small ramp on the magnetic field strength and concomitantly on the resonance frequency of the nuclear spins, along a single axis. Three gradient coils with orthogonal axes are used to “spatially encode” the MR signal by creating a signature resonance frequency at each location in the body. Radio frequency (RF) coils are used to create pulses of RF energy at or near the resonance frequency of the hydrogen nuclei. These coils are used to add energy to the nuclear spin system in a controlled fashion. As the nuclear spins then relax back to their rest energy state, they give up energy in the form of an RF signal. This signal is detected by the MRI system and is transformed into an image using a computer and known reconstruction methods.
MR images may be created by applying currents to the gradient and RF coils according to known methods called “pulse sequences”. A pulse sequence diagram may be used to show the amplitude, phase and timing of the various current pulses applied to the gradient and RF coils for a given pulse sequence. A system operator's selection of a pulse sequence determines the relative appearance of different tissue types in the resultant images, emphasizing or suppressing tissue types as desired. The inherent MR properties of tissue, most commonly the relaxation times T1 and T2, may be exploited to create images with a desirable contrast between different tissues. For example, in a MR image of a brain, gray matter may be caused to appear either darker or lighter than white matter according to the MRI system operator's choice of a T1-weighted or T2-weighted pulse sequence.
Spin Echo (SE) pulse sequences have been used extensively in clinical MR imaging, in part because of their robustness to off-resonance effects such as are caused by main magnetic field inhomogeneity or magnetic susceptibility variations in the imaging subjects. A Spin Echo pulse sequence includes an excitation RF pulse and a single refocusing RF pulse. Typically, the flip angles of the excitation and refocusing pulses are set to 90° and 180° respectively. A spin echo, formed by the refocusing RF pulse, is encoded as a single k-space line and collected at time TE, the “echo time,” after the excitation RF pulse. This combination of pulses and echo acquisitions is repeated at time interval TR, the “repetition time,” until all necessary lines of k-space have been collected. The main advantage of a SE pulse sequence is its ability to create a specific contrast weighting, either T1-, T2-, or proton density-weighted, by modifying the TE and TR.
Fast Spin Echo (FSE) (also known in the art as “Rapid Acquisition Relaxation Enhancement (RARE),” or “Turbo Spin Echo (TSE)”) is a fast version of a SE sequence that uses an excitation RF pulse followed by a train of refocusing RF pulses and resulting spin echoes (i.e., a resulting “echo train”). Typically, the flip angles of the excitation and refocusing RF pulses are set to 90° and 180°, respectively. Multiple spin echoes are collected after each excitation RF pulse, i.e., multiple k-space lines are obtained in a single TR. FSE may be used in either two-dimensional (2D) or three-dimensional (3D) acquisition mode. The maximum practical length of the echo train, i.e., the “Echo Train Length” (ETL), is determined primarily by the T2 relaxation times of the tissues being imaged and the maximum allowable RF power deposition. At 1.5 Tesla, the length of a train of 180° refocusing RF pulses is limited practically to less than approximately 300 ms by the “T2 decay envelope” of the amplitude of the echoes. Typically, TR is significantly longer than TE in order to allow sufficient longitudinal recovery (i.e., “T1 recovery”) of the magnetization after an echo train and before the next excitation. Two-dimensional acquisitions use an interleaving strategy in which k-space data from multiple slices are acquired during a single TR. This acquisition strategy is typically very efficient. Each TR may be partitioned into data acquisition periods for multiple slices, eliminating dead time in the TR. For 3D acquisitions, however, an interleaving strategy is not possible because the entire volume of interest is excited and significant dead time results between the end of an echo train and the next excitation RF pulse.
Refocusing RF pulses with reduced or smaller (i.e., less than 180°) flip angles have been used to reduce RF power deposition and to prolong the amount of time for which magnetization is available for refocusing. The use of reduced flip angles results in the temporary storage of magnetization in stimulated echo coherence pathways during which time the magnetization decays at the tissue's T1 relaxation rate instead of the T2 relaxation rate. Because T1 is significantly longer than T2 for most tissues, the use of refocusing RF pulses with smaller flip angles increases the amount of magnetization available for creating echoes later in the echo train and echo train lengths may be increased while maintaining signal. It is not necessary, however, for all the refocusing RF pulses to have the same flip angle. Rather, a “flip angle schedule” specifying the individual flip angles in a train of RF pulses may be devised to control the RF power deposition and to control aspects of the image appearance such as image contrast, signal-to-noise ratio, blurring, and spatial resolution. For example, instead of using a constant smaller flip angle throughout the RF pulse train, the flip angles of the first refocusing RF pulses may be decreased from a large flip angle to the desired smaller flip angle gradually such that the resultant echoes vary smoothly in magnitude and image artifacts from signal oscillations are eliminated. This preparatory stage in which the flip angles are gradually decreased serves to create a “pseudo-steady state” of the magnetization, in which the resultant echo magnitudes are constant (excluding the effects of relaxation). Flip angles for later refocusing RF pulses in the RF pulse train may be varied to increase the amount of magnetization stored in stimulated echo pathways, thereby maintaining the magnitude of echoes in the echo train and offsetting the effects of T2 relaxation. This allows the use of longer echo trains.
Accordingly, it would be desirable to provide a method and apparatus for calculating or generating a flip angle schedule for long spin echo train pulse sequences with reduced refocusing flip angles. It would be advantageous to provide a method for generating a flip angle schedule that is efficient and that is generally applicable across various tissues and materials (i.e., a method that does not need to be tailored to tissue specific relaxation parameters).